The National Institute of Health (NIH) has concluded that heart failure constitutes “a new epidemic” in the USA. Heart failure, a chronic, progressive and incurable disease, affects over 20 million people worldwide. In the US alone, some 5 million people have been diagnosed with heart failure. Heart failure is estimated to cost the US economy today more than $40 billion annually.
Intracardiac pressure management is an important aspect of heart failure treatment. For example, a rise of the intracardiac pressure, such as in the left atrium is an important early indication of disease progression and the first opportunity for therapeutic intervention. Current blood pressure-measuring methods only can be applied in the coronary care unit (CCU) or the intensive care unit (ICU) and provide no more than an occasional snapshot of intracardiac pressure when the patient is already in a very critical situation. The limitations on current intracardiac pressure measurement methods are a serious impediment to early and optimal treatment. Current treatment methods require hospitalization and may be extremely costly (on average, over $16,000 per patient admittance). The ability to monitor patients and intervene outside of the hospital setting would greatly reduce the number of hospitalizations and extend the lives of those affected by the diagnosis.
Various sensors and devices have been used or proposed for the measurement and analysis of the blood pressure and/or temperature of a patient with mixed success. The currently contemplated sensors have certain disadvantages. For example, the telemetric sensor described in U.S. Pat. No. 6,855,115 can be implanted in the heart by a catheter and is not designed for surgical implantation. Moreover, the sensor, which is rolled up during the implantation procedure, must be made of a flexible material of a specific configuration so that any change of the blood pressure inside the heart effectuates a change in the distance of the sensor height, i.e., the distance between the two capacitor plates used in the sensor. This flexible sensor is folded for delivery via a catheter and then unfolded at the place of implantation. However, a disadvantage of such a configuration is its required flexibility as constant and precise acquisition of measurement data may not be possible when the sensor is placed on or close to the cardiac muscle, and therefore is exposed to the cardiac motions, which may influence correct pressure readings. In addition, the flexible material of a sensor made in accordance with U.S. Pat. No. 6,855,115 may deform due to exposure to constantly streaming liquids, especially a turbulent blood stream likely encountered inside the heart. As a consequence, the capacitance of the capacitor may be changed and measurement values may deteriorate and/or deviate from the true value. Another disadvantage of this type of sensor is due to its use of a pressure-dependent LC-oscillator. The resonant frequency of this oscillator can be analyzed telemetrically. In principle, this kind of device can be applied to measure the pressure that affects the measurement capacitor. Thus, any damage to the material can affect the pressure measurements obtained. Further, as the sensor is influenced by the surrounding media of the sensor, a corruption of measurement values may occur. In addition, there is no circuitry in this type of sensor to digitize the pressure measurement values acquired. Using analog signals may result in external interference during the acquisition and transmission of data, which causes inaccuracies in readings.
Another exemplary implantable device, described in U.S. Pat. No. 6,409,674, uses a catheter filled with a pressure transmitting fluid or gel-like material. The catheter transmits pressure to a pressure transducer within a housing. The sensed pressure is then telemetrically transmitted to an external reader. However, such a device requires a housing for the electronic signal processing circuitry, which results in a larger and heavier sensor structure that can cause strain on the heart when implanted into a heart wall. Moreover, the catheter and housing configuration creates a more complicated, mechanical structure that may be at increased risk for mechanical failure, and therefore is not suitable for long term implantation.
Another device, described in U.S. Pat. No. 6,970,742, has a pressure sensor placed within the heart. A signal from the pressure sensor is transmitted to a housing outside the heart which contains the electronic processing circuits. The signal is processed by the electronic processing circuits, such as converting the signals from analog to digital, and then telemetrically transmitted to an external reader. However, housing the electronic processing circuitry requires additional components and a relatively larger implanted device. Moreover, because digitization of the signal does not occur until outside of the heart, there is a risk of interference in the wire connecting the sensor and the electronic processing circuitry, as analog interference may result from external sources.
Small pressure sensor chips including the electronic processing circuits have been used in other applications. For example, integrated chips having pressure sensors have been used for pressure measurement in optical and cranial applications. These sensors are compact and have fewer mechanical components. Examples of such pressure sensor chips are described in EP 1 312 302 A2 and German patent application DE 10 2004 005 220.7, in which the inventors of the present invention were involved. However, these integrated chips are used in a relatively stable environment, with little movement in the fluids of the eye or brain. Nor are these pressure sensors subject to the cyclical, dynamic movements found in the heart. Such movement may harm connections, such as connections between wires and the pressure sensing chip. Thus, the use of such pressure sensor chips is not suited for the environment of the heart, where there is cyclical and dynamic movement, and where there is continuous and turbulent fluid movement around the pressure sensor.
Conventional techniques to provide stability and support to such known pressure sensing chips to enable their use as a cardiovascular pressure sensors would not likely succeed. Directly attaching a wire to a pressure sensing chip may have a negative impact on the functionality of the chip. For example, when soldering is used for the connection, the heat may damage the chip. One known method of avoiding that problem is to adhere a substrate to the back of the pressure sensing chip, solder the wire to a bond tack on that substrate, and then connect the wire to the chip. However, such substrates have different coefficients of thermal expansion than the chip. Thus, as the temperature changes, the substrate expands and contracts at a different rate then the pressure sensor chip, thereby causing stress and strain on the pressure sensing chip and increasing the risk of damage and/or inoperability.
Other known pressure sensors require a cable connection between the pressure sensor inside the heart and the external body monitoring device. However, such a cable clearly requires an entry into the body. An entry may be inconvenient and require the implantation of both the device and the entry, as well as increase the risk of infection for the patient.
Thus, there is a need for intra-cardiac pressure sensors that are more reliable and accurate, and which cause less irritation when implanted in the heart and are more compatible with the dynamic conditions encountered in a moving heart. Also, a need exists for such a sensor to be used at other locations within the cardiovascular system with little or no modifications.